Field of the Invention
The invention concerns a method and magnetic resonance apparatus for generating a parameter map that, for a target volume in an examination object, represents a field deviation from the resonance frequency of a spectral component of examination tissue of the examination object in a magnetic resonance tomography system (MR system).
Description of the Prior Art
The invention also concerns a method for generating magnetic resonance image data of a target volume in an examination object by the operation of a magnetic resonance tomography system, in which a parameter map generated in this way is used for shimming a B0 field of the magnetic resonance tomography system and/or for determining a current operating frequency of the magnetic resonance tomography system. The invention also concerns a controller for a magnetic resonance tomography system for implementing such a method and to a magnetic resonance tomography system having such a controller.
In a magnetic resonance tomography system, the body to be examined is conventionally exposed in a scanner, having a basic field magnet, to a relatively high basic magnetic field, having a field strength, for example, of 1.5 tesla, 3 tesla or 7 tesla. By the application of the basic field, nuclei in the examination object align according to their nuclear magnetic dipole moment, frequently also called spin, along the field. This collective behavior of the spin system is described as macroscopic magnetization. The macroscopic magnetization is the vector sum of all microscopic magnetic moments in the object at a specific location. In addition to the basic field a magnetic field gradient is applied by a gradient system. The magnetic resonance frequency (Larmor frequency) applicable at the respective location is directly proportional to the total magnetic field (known as the B0 field) that is present at the respective location due to the superimposition of the basic magnetic field and the gradient magnetic field. Radio-frequency excitation signals (RF pulses) are then emitted by a radio-frequency transmitting system by means of suitable antenna devices, and this leads to the nuclear spins of specific nuclei that are resonantly excited (i.e. at the Larmor-frequency present at the respective location) by this radio-frequency field being tilted by a defined flip angle with respect to the magnetic field lines of the basic magnetic field. If an RF pulse of this kind acts on spins that have already been excited, then these can be tilted into a different angular position or even be deflected back into an initial state parallel to the basic magnetic field. When the excited nuclear spins are relaxed, radio-frequency signals, known as magnetic resonance signals, are resonantly emitted, and these are received by suitable receiving antennae. After demodulation and digitization and possibly further processing steps, the received signals are in the form of complex numbers, called raw data. The magnetic resonance signals are acquired in the spatial frequency domain, known as k-space, and the raw data are entered into a memory representing in a timed k-space sequence of data entry points along a “gradient trajectory” (also called “k-space trajectory”) defined by the switching of the gradient pulses during a measurement. The RF pulses must be emitted so as to be appropriately coordinated in time with the gradient activations. The desired image data (MR images) can be reconstructed from the raw data acquired in this way. This image reconstruction frequently includes a two-dimensional Fourier transformation.
Specific, predefined pulse sequences are conventionally used to activate a magnetic resonance tomography system during the measurement, i.e. sequences of defined RF pulses and gradient pulses in different directions, and readout windows, during which the receiving antennae are switched to a receive state and the magnetic resonance signals are thus received and processed. With the use of what is known as a measurement protocol, these sequences are parameterized in advance for a desired examination, for example to give a specific contrast to the calculated images. The measurement protocol can also include further control data for the measurement. There are many magnetic resonance sequence techniques according to which pulse sequences can be established.
Many magnetic resonance technologies or measuring methods, such as spectral fat suppression, or fast imaging methods, such as EPI (Echo Planar Imaging) or spiral techniques, make high demands on the homogeneity of the B0 field. Each individual body of each patient disturbs the local magnetic field differently. To nevertheless be able to use such methods, in practice a technique known as in-vivo shimming (a field adjustment with the patient positioned in the device) is often carried out patient-specifically. In this shimming, first the local B0 field (i.e. the field present at the respective image point) in the examination region is measured to create a record known as a B0 map, as mentioned in the introduction. DC offset currents are then calculated from the B0 map for the three gradient coils (i.e. the linear shim terms or terms of the field deviation) as well as currents for specific shim channels (or shim coils) of a higher order which compensate the local field distortions in the best possible way. After adjusting these currents an RF resonance frequency is ascertained as a rule in a frequency adjustment for the desired spectral component of the examined tissue (usually protons bound to water) which are then specified to the components of the system, in particular the RF transmitting system and the RF receiving system as the operating frequency in order to emit RF pulses with the appropriate carrier frequency f0 and receive magnetic resonance signals.
In the case of a measurement known from practice for adjusting a B0 map, two complex MR images are measured at different echo times T1 and T2 for example with a double echo gradient sequence or DESS (Double Echo Steady State) sequence to calculate the local off-resonance frequency Δf(x,y,z) (i.e. the deviation from the resonance frequency) from the phase of the difference image at location (x,y,z) (i.e. a phase difference map ΔΦH(x,y,z)):
                              Δ          ⁢                                          ⁢                      f            ⁡                          (                              x                ,                y                ,                z                            )                                      =                              ΔΦ            ⁡                          (                              x                ,                y                ,                z                            )                                            2            ⁢                          π              ⁡                              (                                                      T                    2                                    -                                      T                    1                                                  )                                                                        (        1        )            
This method is based on the assumption that the phase accumulation between the two echo times is solely a consequence of the local deviation from the RF resonance frequency (RF center frequency). With a presence of a plurality of spectral components (i.e. components in examination tissues with different resonance frequencies) this assumption is only correct if the relative phase position of the relevant spectral components does not change between the two echo times. In the case of just two dominant spectral components this may be achieved by choosing the echo time difference in such a way that the phase evolution of one component is a multiple of 2π compared with the other component between the two echo times. The two dominant spectral components in the proton imaging most used in practice are protons bound to water and fat. Their resonance frequencies are shifted by roughly 3.2 to 3.4 ppm (“parts per million”) with respect to each other, in a magnetic field of 1.5 T by approx. Δfc=204 Hz therefore and at 3 T by approx. Δfc=408 Hz (“c” stands for “chemical shift”). This in turn corresponds to a minimum echo time difference of 4.86 ms at 1.5 T and 2.43 ms at 3 T.
As a consequence of the 2π periodicity of the arc tan 2 function, with the aid of which the phase can be determined from the complex difference image, this choice of the echo times according to equation (1) unavoidably limits the maximum off-resonance frequency Δf(x,y,z), which can be clearly determined using this method, to ±Δfc/2, i.e. to approx. ±102 Hz at 1.5 T and ±204 Hz at 3 T. Higher off-resonances (i.e. deviations from the resonance frequency) lead to what are known as phase wraps in the calculated B0 maps.
The off-resonances that actually occur are usually larger. The B0 maps determined using the method therefore exhibit phase wraps as a rule. In practice, only the DC offset currents are therefore currently ascertained from the B0 maps for the gradient coils and the higher shim currents. Methods are known in practice for this purpose which function robustly even in the presence of phase wraps. The new resonance frequency is in practice usually ascertained following the adjustment of these shim currents using a spectroscopic method, i.e. the Fourier analysis of an MR signal, which is received without simultaneous switching of gradients. This separate adjustment step is also called a “frequency adjustment”. Generally only one frequency for the entire measuring volume is ascertained. Due to the long repetition time of the sequences used for the spectroscopic method, such as STEAM (“stimulated echo acquisition mode”), a spatially resolved frequency adjustment would be very time-consuming.
The article that appeared in the journal MRM 38 on pages 477-483 1997 “Dynamic shimming for Multi-Slice Magnetic Resonance Imaging” by Glen Morrell and Daniel Spielman proposes measuring a B0 map using a double echo-gradient echo sequence. Fat-water errors are prevented by appropriate choice of the echo time difference and the spectral range of the B0 map is expanded by phase unwrapping (“phase unwrapping”). However, the B0 map then always still contains an unknown global offset.
A further example of a double echo method in the prior art is the ISMRM 2011 Abstract by Dan Xu et al. with program number 2689. There it is proposed to determine the local frequency from two reference scans which are acquired in echo planar-imaging for phase correction. Xu et al do not discuss the problems that result due to the presence of various spectral components in the examination tissues.